KR20200019040A - Biosensor surface-modified with neutral agent and detecting method using the same - Google Patents

Abstract

The present invention relates to a biosensor whose surface is modified with a neutral detergent, and a detection method using the same. According to the sensor according to one aspect, it is possible to more effectively block the formation of a non-specific binding of an unknown substance in a sample on the surface of the sensor, thereby maximizing detection efficiency. In addition, the detection method may not involve a washing step, thereby providing an effect of making an assay procedure simpler and faster.

Description

Biosensor surface-modified with neutral agent and detection method using the same

It relates to a biosensor surface-modified with a neutral formulation and a detection method using the same.

Biosensors, which measure the presence or absence of certain biological materials, are devices that convert physical or chemical changes caused by selective interactions between biological elements and analytes into recognizable optical or electrical signals. . Applications of such biosensors are for use in the analysis of pollutants in the environment, for the detection of biochemical weapons for mass destruction in the military, and for the detection of harmful or decaying substances in the food sector. However, especially in the clinical diagnosis and medical field has been spotlighted for the purpose of early diagnosis of diseases or analysis of biological material.

However, as a limitation of biosensors, biological elements that specifically bind to analyte, such as unknown substances contained in a sample in a region other than the region where the antibody is immobilized, may cause nonspecific interaction with the sensor surface. Noise signals may occur.

Field-effect transistor-based biosensor according to an embodiment is highly sensitive to the external environment because it is a high sensitivity sensor. In particular, the surface treatment for stable signal detection is important because it is very sensitive to surface charge. The conventional blocking method, bovine serum albumin (BSA) coating, has a large negative charge, which can induce nonspecific binding to various proteins contained in serum, which may cause noise in the detection signal. Patent 10-2018-0043916). Therefore, there is a need for a method of controlling surface charge by introducing a new blocking method.

One aspect includes a reaction unit in which physical or chemical interaction occurs with a target material in a sample; And a signal converter configured to recognize the interaction and convert the interaction into an electrical signal. An electrode formed on the substrate; It is to provide a biosensor comprising a blocking layer comprising an electrically neutral formulation formed on the electrode.

Another aspect includes contacting a sample with the biosensor; And it provides a target material detection method comprising the step of observing the electrical conductivity change of the biosensor.

Another aspect is a substrate; electrode; Providing a reaction unit including a test cell; Treating the electrode surface using a linker; Immobilizing a biological probe to the linker; And it provides a method for producing a biosensor comprising the step of forming a blocking layer on the electrode with an electrically neutral formulation.

One aspect includes a reaction unit in which physical or chemical interaction occurs with a target material in a sample; And it provides a biosensor comprising a signal converter for recognizing the interaction and converting into an electrical signal.

The sample may be a biological sample derived from an individual, eg, a mammal, including a human. In addition, the biological sample may be blood, whole blood, serum, plasma, lymph, urine, feces, tissues, cells, organs, bone marrow, saliva, sputum, cerebrospinal fluid or a combination thereof.

Physical or chemical interactions with the target material are not particularly limited, but are usually caused by the forces acting between molecules resulting from at least one of covalent bonds, hydrophobic bonds, hydrogen bonds, van der Waals bonds, and electrostatic bonds. Indicates. Covalent bonds include coordination bonds and dipole bonds. In addition, coupling by electrostatic power includes electric repulsion in addition to electrostatic coupling. Moreover, the coupling | bonding reaction, the synthetic | combination reaction, and the decomposition reaction resulting from the said action are also contained in interaction. Specific examples of interactions include binding and dissociation between antigen and antibody, binding and dissociation between protein receptor and ligand, binding and dissociation between adhesion molecule and partner molecule, binding and dissociation between enzyme and substrate, binding between apo-enzyme and coenzyme, and Dissociation, binding and dissociation between nucleic acids and proteins that bind to them, binding and dissociation between nucleic acids and nucleic acids, binding and dissociation between proteins in the information delivery system, binding and dissociation between sugar proteins and proteins, or binding and dissociation between sugar chains and proteins , Binding and dissociation between cells and biological tissues and proteins, binding and dissociation between cells and biological tissues and small molecule compounds, and interactions between ions and ion-sensitive substances.

In one embodiment, the reaction unit; An electrode formed on the substrate; It may include a blocking layer comprising an electrically neutral formulation formed on the electrode. The reaction unit may be configured to be used for single use. For example, the substrate may be a material selected from the group consisting of silicon, glass, metal, plastic, and ceramic. Specifically, the substrate may be selected from the group consisting of silicon, glass, polystyrene, polymethylacrylate, polycarbonate, and ceramic. Examples of the electrode may be titanium nitride, silver, silver epoxy, palladium, copper, gold, platinum, silver / silver chloride, silver / silver ions, or mercury / mercury oxide. In addition, the reaction unit may include an insulating electrode formed on the substrate or the electrode. The insulating electrode may include an oxide film formed naturally or artificially. Examples of the oxide film may include SixOy, HxfOy, AlxOy, TaxOy, or TixOy (where x or y are integers of 1 to 5). Forming the oxide film can be made by a known method. For example, the oxide may be formed by depositing liquid phase deposition, evaporation, and sputtering on a substrate. A test cell for accommodating a sample may be attached onto the insulating electrode. The test cell is polydimethylsiloxane (PDMS), polyethersulfone (PES), poly (3,4-ethylenedioxythiophene) (poly (3,4-ethylenedioxythiophene)), poly (styrenesulfonate) (poly (styrenesulfonate)), polyimide, polyurethane, polyester, perfluoropolyether (PFPE), polycarbonate, or a combination of the above polymers It may be. As used herein, the term "blocking layer" may refer to a layer obtained by applying the surface of the electrode using one or more blocking agents.

In one embodiment, the electrically neutral agent may be a neutral protein or neutral peptide. The electrically neutral formulation may mean a substance in which the net charge amount of the molecule becomes zero when the intrinsic isoelectric point (pI) of the substance corresponds to 6.3 to 7.5 and the pH of the solvent is 6.3 to 7.5. Any material that makes the surface charge neutral is possible. For example, in an amphoteric electrolyte containing both anionic and cationic groups, such as proteins and peptides, this may mean that the net charge of the molecule becomes zero.

In one embodiment, the neutral peptide is a polymer composed of any amino acid combination, it may be a material having a pI value of 6.3 to 7.5 of the entire peptide. Accordingly, the amino acid may be a negatively or positively charged amino acid because the pI value does not correspond to 6.3 to 7.5. In addition, the peptide may be composed of 10 to 100 amino acids.

In one embodiment, the neutral protein is, for example, myogen (pI = 6.3), gliadin (pI = 6.5), hemoglobin (pI = 6.8), and myoglobin (mioglobin, pI). = 7.0) or combinations thereof. The net charge of the protein may be determined by amino acids on the protein surface rather than amino acids throughout the protein.

The biosensor may recognize any electrochemical sensor that is highly influenced by surface charge in recognizing interaction between a biological element and an analyte, and according to an embodiment, a field-effect transistor (FET) It may be based on a biosensor.

In one embodiment, the signal converter may include a field effect transistor. The signal conversion unit may be detachable from the reaction unit, and may be made of a connection between an electrode of the reaction unit and an upper gate electrode of the transistor. The connection can take the form of a plug, for example.

In one embodiment, the reaction unit may include a biological probe specifically binding to the target material.

As used herein, the term “biological probe” may refer to a substance that can impart functionalization to a reaction part or a substance that specifically binds to a target substance. The biological probe may be DNA, RNA, nucleotides, nucleosides, proteins, polypeptides, peptides, amino acids, carbohydrates, enzymes, antibodies, antigens, receptors, viruses, substrates, ligands and membranes or combinations thereof. For example, it may be ANXA 3 antibody that specifically binds to ANXA 3 antigen. In addition, the biological probe may include a redox enzyme. The redox enzyme may refer to an enzyme for oxidizing or reducing a substrate, and may include, for example, an oxidase, peroxidase, reductase, catalase, or dehydrogenase. Examples of the redox enzymes include blood sugar oxidase, lactate oxidase, cholesterol oxidase, glutamate oxidase, horseradish peroxidase, alcohol oxidase, glucose oxidase (GOx), glucose dehydrogenase (GDH), Cholesterol estergenase, ascorbic acid oxidase, alcohol dehydrogenase, laccase, tyrosinase, galactose oxidase or bilirubin oxidase have.

In one embodiment, the reaction unit may include a linker for immobilizing the biological probe on the surface of the electrode. As used herein, the term “immobilization” may mean that a biological probe forms a chemical or physical bond to a substrate or electrode. The linker is biotin, avidin, streptavidin, carbohydrate, poly L-lysine, hydroxyl group, thiol group, amine group, alcohol group, carboxyl group, amino group, sulfur group, aldehyde group, carbonyl group, succinimide group, maleimide group, epoxy group And compounds having isothiocyanate groups or combinations thereof. As a specific example, it may be a combination of APTES and glutaraldehyde.

As used herein, the term “target material” may refer to a material that can be present in a sample and specifically refers to a material that specifically binds to a biological probe. Detectable targets may include those that may be involved in specific binding interactions with one or more biological probes that may participate in a sandwich, competition, or substitution assay configuration. Examples of such targets include antigens such as peptides (eg hormones), hapten, carbohydrates, proteins (eg enzymes), drugs, pesticides, microorganisms, antibodies, and complementary sequences with sequence specific hybridization reactions. Nucleic acids that can participate in or combinations thereof. As a specific example, it may be ANXA 3 protein.

In one embodiment, the electrically neutral agent may be to block the target material from forming a nonspecific bond in a region other than a region that forms a specific bond with the biological probe.

As used herein, the term “nonspecific binding” may mean that an unknown material included in a sample causes a noise signal by performing nonspecific interaction with the sensor surface, such as hydrophobic interaction and / or charge interaction.

In the reaction section, a sample enters through the test cell for receiving the electrode, the biological probe and the target material, and the target material present in the sample combines with the biological probe to cause a chemical potential gradient in the test cell. The term "chemical potential gradient" may refer to the concentration gradient of the active species. When such a slope is present between the two electrodes, the potential difference may be detected when the circuit is opened and the current will flow until the slope is lost when the circuit is closed. The chemical potential gradient may refer to any potential gradient resulting from the application of a potential difference or current flow between the electrodes.

In one embodiment, the field effect transistor comprises a substrate; Insulating layer; Source and drain electrodes spaced apart from each other; A gate electrode; It may include a channel layer disposed between the source electrode and the drain electrode. In an embodiment, the field effect transistor includes: a lower gate electrode; A lower insulating film formed on the lower gate electrode; A source and a drain formed on the lower insulating film and spaced apart from each other; A channel layer formed on the lower insulating film and disposed between the source and the drain; And a dual gate field effect transistor including an upper insulating layer formed on the source, the drain, and the channel layer, and an upper gate electrode formed on the upper insulating layer.

The small surface potential voltage difference generated in the reaction unit greatly amplifies the threshold voltage change of the lower field transistor due to the supercapacitive coupling generated in the double gate ion sensing field effect transistor (ISFET) including the channel layer. The amplification factor may be determined by the thickness of the lower insulating film, the thickness of the channel layer, and the thickness of the insulating film of the upper gate. The thicker the lower insulating film, the thinner the upper insulating film and the channel layer, the larger the size of the amplification factor.

The channel layer may be an ultra-thin layer, for example, the thickness may be 10 nm or less, 9 nm or less, 8 nm or less, 7 nm or less, 6 nm or less, 5 nm or less, or 4 nm or less.

Within the range of the thickness of the channel layer, due to the strong electric field of the lower gate electrode that is induced in the ultra-thin film, supercapacitive bonding that can be controlled under all conditions up to the upper interface occurs. Through this, the electrons and holes induced in the upper gate interface may also be controlled and the leakage current may be blocked. In addition, by allowing a stable amplification factor, it is possible to improve the linear response, hysteresis, and drift phenomenon according to the surface potential, and to maintain the electrostatic coupling of the upper and lower gates. In addition, within the range of the thickness of the channel layer, the transistor including the ultra-thin channel layer allows a large amplification factor compared to the conventional transistor, while also increasing the ion sensing power. In addition, the transistor including the ultra-thin channel layer within the range of the thickness of the channel layer can improve the stability compared to the conventional transistor. The varying amplification factor seen in the thick channel layer, coupled with the leakage current component induced at the upper interface, can cause device degradation due to ion damage. On the other hand, the transistor according to the embodiment in which the leakage current is controlled while allowing a constant amplification factor can minimize the ion damage effect. In addition, when the lower insulating layer is excessively thick in the existing transistor, a phenomenon occurs in which the lower electric field does not control all of the channel regions, thereby weakening the electrostatic coupling of the upper and lower gates, which includes an ultra-thin channel layer according to an embodiment. Transistors can obtain large amplification factors while maintaining electrostatic coupling. The electrostatic coupling of the upper and lower gates occurs when the upper channel interface is completely depleted. In the conventional transistor, the amplification does not occur because the entire length of the lower gate does not control the upper channel.

The channel layer may include any one selected from the group consisting of an oxide semiconductor, an organic semiconductor, polycrystalline silicon, and single crystal silicon. When the channel layer includes any one selected from the group consisting of semiconductors, organic semiconductors, polycrystalline silicon, and single crystal silicon, upper and lower gate electrostatic coupling occurs and high sensitivity sensors can be manufactured, and transparent and flexible sensors can be provided. have. The channel layer is not limited in width or length, and may use an electrostatic coupling phenomenon by using upper and lower gate electrodes in a double gate structure.

In the sensor, an equivalent oxide thickness of the upper insulating layer may be thinner than an equivalent oxide thickness of the lower insulating layer. For example, the thickness of the upper insulating film may be about 25 nm or less, and the thickness of the lower insulating film may be about 50 nm or more. When the equivalent oxide film thickness of the upper insulating film is thinner than the equivalent oxide film thickness of the lower insulating film, a sensitivity amplification phenomenon may be caused.

The upper insulating film and the lower insulating film may include an oxide film formed naturally or artificially. Examples of the oxide film may include Si _x O _y , H _x fO _y , Al _x O _y , Ta _x O _y , or Ti _x O _y (where x or y are integers of 1 to 5). The oxide film may have a single, double, or triple stacked structure. Through this, by increasing the physical thickness, and by reducing the equivalent oxide film thickness of the upper insulating film, it is possible to amplify the sensitivity of the sensor, and to prevent the phenomenon of degradation due to leakage current.

The double gate ion sensing field effect transistor according to an embodiment may have a structure including a field transistor including an upper insulating layer and a lower field transistor including a lower insulating layer in one device. According to each operation mode, it can operate independently with the upper and lower gates. When the upper and lower gates of the device are used at the same time, while the electrostatic coupling phenomenon is observed due to the structural specificity of the double gate structure, the interconnection of the upper and lower field transistors can be established. The dual operation mode may be to use the lower gate as the main gate. Thus, the transistor according to one embodiment may be operating in the double gate mode.

The sensor may include a plurality of reaction units and a plurality of transistors to detect a plurality of target materials. The sensor may include a plurality of reaction units and a plurality of ion sensing field effect transistors, and the plurality of reaction units and the plurality of ion sensing field effect transistors may be electrically connected to each other. A plurality of sources in the plurality of transistors may be commonly grounded, a plurality of upper gate electrodes may be commonly grounded, and a plurality of lower gate electrodes may be applied with a common voltage. For example, the source of the first transistor and the second transistor, and the reference electrodes of the first and second reaction units may be commonly grounded. For example, a constant common voltage may be applied to the lower electrodes of the first transistor and the second transistor. In addition, the plurality of drains in the plurality of transistors may have a parallel structure. For example, the drains of the first transistor and the second transistor may have a parallel structure. In addition, the plurality of reaction units may be one in which different biological probes are immobilized. The plurality of transistors may detect the same or different target material signals from the plurality of reaction units, amplify them, and output a signal through a semiconductor parameter analyzer.

The signal converter may be electrically connected to the transistor, and may further include a calculation module for determining an amount of a target material in a sample from a potential difference measured by the transistor. The calculation module may be for determination of a target material. As used herein, the term "determining a target substance" may mean a qualitative, semi-quantitative and quantitative process for evaluating a sample. In qualitative evaluation, the results indicate whether the target material is detected in the sample. In semi-quantitative evaluations, the results indicate whether the target material is above a predefined boundary value. In quantitative evaluation, the result is a numerical representation of the amount of target material present. Conversion of measured values can also use a look-up table that converts specific values of currents or potentials to values of the target material depending on the specific device structure and calibration values for the target material. The calculation module can be determined by measuring the potential difference according to a known concentration of the target material. For example, the calculation module may be to determine the amount of the target material in the sample compared to the normal control.

Another aspect includes (a) contacting a sample with the biosensor; And (b) provides a target material detection method comprising the step of observing the electrical conductivity change of the biosensor.

As used herein, the term “detection” may mean detecting or identifying the presence or presence of a target substance, and may include, for example, identifying the target substance or quantifying the target substance in a sample.

The detection method may be an electrochemical method of measuring the current or potential generated by the target material interacting with the biological probe.

In one embodiment, the step (b) may be performed after the step (a) without the washing step. For example, in order to remove the noise signal due to nonspecific binding caused by passing through a serum sample, washing with PBS is omitted, and the change of electrical conductivity of the biosensor in the serum state is omitted immediately. Can be observed.

According to the sensor according to one aspect, it is possible to more effectively block the formation of non-specific binding of unknown substances in the sample on the surface of the sensor to maximize the detection efficiency and may not be accompanied by a washing step, so that There is an effect that can be done more quickly.

1 measures the Vg voltage shift of the IV curve in serum prior to washing.
Figure 2 is a re-measurement of the Vg shift value of the I-V curve after washing with PBS.

Hereinafter, the present invention will be described in more detail with reference to Examples. However, these examples are for illustrative purposes only and the scope of the present invention is not limited to these examples.

Example 1 Fabrication of Surface-Modified FET-Based Biosensors

Example 1-1. Fabrication of Double Gate Ion Sensing Field Effect Transistors

The substrate is made of silicon-on-insulator (SOI) having a resistivity of about 10 to 20 Ωcm, the thickness of silicon, the lower gate electrode, is about 107 nm, and the thickness of the buried SiO2 oxide film, the lower insulating film, is about 224 nm. It was. After performing standard RCA cleaning, the upper silicon was etched with about 2.38% by weight of tetramethylammonium hydroxide (TMAH) solution to form an ultra thin film, and a channel region was formed using a photolithography graph. The length and width of the formed channels were about 20 um and 20 um, respectively, and the thickness was about 4.3 nm. A titanium nitride (TiN) electrode was then formed using a sputtering system. Thereafter, about 23 nm thick silicon dioxide was formed on the source and drain through oxidation to form an upper insulating film. The upper gate electrode was deposited using a sputtering system on a TiN thin film layer having a thickness of about 150 nm. Next, a double gate ion sensing field effect transistor was manufactured by performing heat treatment at a gas condition including N 2 and H 2 at a temperature of about 450 ° C. in order to eliminate defects and improve the interface state therebetween.

Example 1-2. Manufacture of reaction part

In order to fabricate the reaction part, a substrate of about 300 nm was used. After standard RCA cleaning, an E-beam evaporator was used to deposit a working electrode ITO at about 100 nm thick to measure the electrical potential difference on the substrate surface. Next, as an insulating electrode, a SnO 2 film, which is an oxide film, was deposited to a thickness of about 45 nm on the ITO layer using an RF sputter. At this time, RF power was about 50 W. Thereafter, the sputtering process was performed under Ar gas conditions and about 3 mtorr pressure conditions with a flow rate of about 20 sccm. Subsequently, a test cell for accommodating a sample was made of polydimethylsiloxane (PDMS) and attached on the insulating electrode to prepare a reaction part. The test cell was fabricated with an extended gate (EG) with six wells. In addition, a silver / silver chloride electrode was used as a reference electrode. After surface treatment with 5% APTES and 2.5% glutaraldehyde, the ANXA3 antibody was immobilized on the surface by reacting at room temperature for 1 hour. 1% of BSA (-), hemoglobin (neutral), and lysozyme (+) buffer dissolved in PBS were reacted at room temperature for 1 hour, and then washed three times with PBS. The six wells consisted of three wells with no antibody attached to the surface and the remaining three wells with antibody attached to the surface. The BSA (-), hemoglobin (neutral), and lysozyme (+) buffer were reacted in wells in which the antibody did not adhere to the surface and in wells in which the antibody was attached to the surface. The molecular weight, isoelectric point, and charge of the blocking protein BSA, hemoglobin, lysozyme and the target protein ANXA3, and the solvent at pH 7.4 are shown in Table 1 below.

Blocking and Target Protein M.W. (kDa) pI Charge (at pH 7.4) Bovine serum albumin (BSA) 66.3 4.8 Anionic (-) Hemoglobin (Hemo) 64.5 6.8 Neutral Lysozyme (Lyso) 14.3 11 Cationic (+) ANXA3 (target protein) 36 5.6 Anionic (-)

Example 1-3. Fabrication of the sensor

The sensor was manufactured by connecting the upper gate electrode of the transistor manufactured in (1-1) and the working electrode of the reaction part manufactured in (1-2) in the form of a plug-in.

Experimental Example 1. Comparison of detection efficiency before and after washing according to blocking protein

In order to confirm the detection efficiency according to the type of blocking protein, the voltage shift of the I-V curve was measured using ANXA3 protein as a model target. A solution in which 1 ng / ml of ANXA3 antigen was dispensed into 100% serum in 6 wells was reacted at room temperature for 20 minutes. After passing through the serum sample, the threshold voltage change of the gate voltage due to charge transfer that occurs when the biomaterial is adsorbed on the surface of the initial I-V curve (I: current, Vg: gate voltage) was measured. First, the signal was measured immediately in the serum state before washing, and after washing with PBS again to measure the detection efficiency.

Figure 1 measures the Vg shift value of the I-V curve in serum prior to washing. 2 is a remeasurement of the Vg shift value of the I-V curve after washing with PBS.

As a result, as shown in FIG. 2, when hemoglobin, a neutral blocking protein, was used, the difference in voltage shift between the well with no antibody and the well with antibody was found to be greatest. . This indicates that blocking the surface with neutral proteins can minimize nonspecific binding to other proteins with negative charges in the serum in areas other than the region where the antibody is immobilized. In particular, unlike the conventional enzyme-linked immunosorbent assay (ELISA) in which the primary antibody is immobilized and then the secondary antibody is combined with a luminescent factor and converted into an optical signal, the antibody depends on whether the antibody binds to the antigen. In field effect transistor-based biosensors that convert the difference to an electrical signal, the charge of the blocking protein is important because charge-charge interaction is a more important factor in detection efficiency. Therefore, it can be seen that when the surface modification to the neutral, it is possible to maximize the detection efficiency and blocking efficiency in the field-effect transistor-based biosensor.

In addition, as shown in FIGS. 1 and 2, when blocked with hemoglobin, which is a neutral protein, the antibody is attached to the well (no Ab) to which the antibody is not attached even in the serum state before washing, as after washing with PBS. It was confirmed that the voltage shift value appeared larger in the well. On the other hand, in the case of BSA or lysozyme, unlike washing after PBS, in the serum state before washing, the voltage shift value in the no-ab antibody-free well is larger than that in the antibody-attached well. It confirmed the point which appears. This means that a washing step is necessary for BSA or lysozyme, but may not involve a washing step if blocked with a neutral protein such as hemoglobin. Thus, it can be seen that the assay procedure can be made faster if the surface modification is neutral.

Claims (14)

A reaction part in which physical or chemical interaction occurs with the target substance in the sample; And a signal converter for recognizing the interaction and converting the signal into an electrical signal.
The reaction unit substrate; An electrode formed on the substrate; A biosensor comprising a blocking layer comprising an electrically neutral agent formed on the electrode.
The biosensor of claim 1, wherein the electrically neutral agent is a neutral peptide or neutral protein.
The method of claim 2, wherein the neutral peptide is pI (isoelectric point) value of 6 to 8, the biosensor.
The biosensor of claim 2, wherein the neutral protein is at least one selected from the group consisting of hemoglobin, myogen, gliadin, and myoglobin.
The biosensor of claim 1, wherein the signal converter comprises a field effect transistor.
The biosensor of claim 1, wherein the reaction part comprises a biological probe specifically binding to the target material.
The method of claim 6, wherein the biological probe is selected from the group consisting of DNA, RNA, nucleotides, nucleosides, proteins, polypeptides, peptides, amino acids, carbohydrates, enzymes, antibodies, antigens, receptors, viruses, substrates, ligands and membranes. The biosensor which is 1 or more.
The biosensor of claim 6, wherein the reaction part comprises a linker for immobilizing the biological probe on the surface of the electrode.
The method according to claim 8, wherein the linker is biotin, avidin, streptavidin, carbohydrates, poly L- lysine, hydroxyl group, thiol group, amine group, alcohol group, carboxyl group, amino group, sulfur group, aldehyde group, carbonyl group, succinimide group, The biosensor that is at least one selected from the group consisting of a compound having a maleimide group, an epoxy group, and an isothiocyanate group.
The method according to claim 1, wherein the target material is one or more selected from the group consisting of antigens, hapten, carbohydrates, proteins, drugs, pesticides, microorganisms, antibodies, and nucleic acids capable of participating in sequence-specific hybridization reactions with complementary sequences , Biosensor.
The biosensor of claim 1, wherein the electrically neutral agent prevents the target material from forming a nonspecific bond in a region other than a region that forms a specific bond with a biological probe.
The device of claim 5, wherein the field effect transistor comprises: a substrate; Insulating layer; Source and drain electrodes spaced apart from each other; A gate electrode; And a channel layer disposed between the source electrode and the drain electrode.
(a) contacting a sample with a biosensor according to any one of claims 1 to 12; And (b) observing a change in electrical conductivity of the biosensor.
The method of claim 13, wherein after step (a), step (b) is performed without a washing step.

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